Device for generating images and/or projections

ABSTRACT

The invention relates to a device for generating images and/or projections, which device includes a device for the detection of input radiation. The device for the detection of input radiation comprises a sensor with a Pr 3+ -activated scintillator for converting the input radiation into UV radiation. The Pr 3+ -activated scintillators have short excitation and decay times.

The invention relates to a device for generating images and/orprojections by means of an imaging method, which device includes adevice for the detection of input radiation which includes at least oneacquisition element which comprises a sensor for converting the inputradiation and a photodiode which converts an optical signal into anelectrical signal. The invention also relates to a device for thedetection of input radiation.

Tomography is a slice imaging method used in the field of medical X-raydiagnosis. According to this method a coupled motion of the tube and thefilm takes place in opposite directions while the patient remainsstationary. A given, selected depth zone is thus sharply imaged on thefilm, whereas parts of the objects which are situated higher or deeperare blurred because of the continuously changing projection.

In X-ray tomography the attenuation of X-rays by a given slice of thebody of the patient is measured by means of a plurality of detectors indifferent projection directions. To this end, a narrow X-ray beam (fanbeam) is formed by means of an X-ray tube and diaphragms. This beamtraverses the relevant part of the body and is attenuated to a differentdegree by the various structures within the body (for example, the skin,fat, muscles, organs, bones).

Exactly opposite the X-ray tube there is arranged a plurality ofdetectors which receive the attenuated signal and process itelectronically so as to propagate it to a computer for evaluation.Subsequently, the X-ray tube and the oppositely situated detectors arerotated slightly further around the patient.

The described procedure is then repeated. Various views (projections) ofthe same slice are thus generated and converted into a grey scale imagein the computer. This image can be viewed and evaluated on a displayscreen or an X-ray film. This technique yields images which containsignificantly more contrast than those produced by a conventional X-raytechnique.

A further method based on the principle of tomography is called EmissionComputer Tomography (ECT) which is based on the measurement in slices oftemporarily incorporated radionuclides (scintigraphy). The emission ofpositrons from ¹⁵O carbon dioxide or ⁶⁸Ge (Positron Emission ComputerTomography or PET) or photons from ^(99m)Tc or ¹²³I (Single PhotonEmission Computer Tomography or SPECT) is then measured. Such nuclearmedical tomography methods offer the advantage that they present theviewer of the tomographically generated image with information whichgoes beyond pure morphology and which possibly also images physiologicalevents.

In contemporary computer tomography two basic types of radiationdetectors can be distinguished: direct converters (for example, xenongas detectors) and scintillation detectors. In the case of solid-statedetectors, made of a scintillation material (scintillator), the visiblelight emitted after excitation is collected by photodiodes. In X-raycomputer tomography solid-state detectors contain either cadmiumtungstate (CdWO₄) or materials on the basis of rare earths. Frequentlyused scintillators are Ce³⁺-doped materials such as, for example,Lu₂SiO₅:Ce or Gd₂SiO₅:Ce. In PET methods or in SPECT methods thescintillator often used is NaI:Tl or bismuth germanate-Bi₄Ge₃ O₁₂ (BGO).Detectors provided with a scintillator in the form of Ce³⁺-dopedmaterials are known, for example, from EP 1 004 899.

The scintillator must satisfy some conditions. For example, thescintillator must have a high density, a high luminous efficiency and ashort excitation and decay time.

The decay time τ of NaI:Tl amounts to 250 ns, that of BGO to 300 ns,that of Lu₂SiO₅:Ce to 40 ns and that of Gd₂SiO₅:Ce to 56 ns. For manyapplications it is desirable to utilize scintillators having evenshorter decay times.

Therefore, it is an object of the invention to provide a device fortomography which comprises a scintillator having a shorter decay time τ.

This object is achieved by means of a device for generating imagesand/or projections by means of an imaging method, which device includesa device for the detection of input radiation which includes at leastone acquisition element which comprises a sensor with a Pr³⁺-activatedscintillator for converting the input radiation into UV radiation and aphotodiode which converts an optical signal into an electrical signal.

Pr³⁺-activated scintillators have short decay times τ in the range below25 ns which, therefore, are even shorter than those of Ce³⁺-activatedscintillators.

Because of the short decay time τ, the integration time can be reducedduring the determination of the intensity of the input radiation, sothat the image rate for the generating of images and/or projections canbe significantly increased. Because of the increased image rate, theoccurrence of artifacts, for example, shadow images, is reduced.Furthermore, the examination time is reduced for the patient, becausemore single images can be measured within a shorter period of time.

The advantageously selected Pr³⁺-activated scintillators in conformitywith claim 2 have short excitation times and short decay times τ.Furthermore, they emit UV radiation in response to excitation by meansof X-rays or γ quanta.

For the advantageously chosen imaging methods in conformity with theclaims 3 to 5 it is important that the scintillator has a short decaytime and hence enables a high image rate. The time window is of majorimportance notably for the PET method and it is particularlyadvantageous to utilize a scintillator having a short decay time in thedevice for determining input radiation.

The advantageous embodiments in conformity with the claims 6 to 8 makeit possible to use a larger part of the input radiation for the imageanalysis. The luminous substance that can be excited by ultravioletradiation absorbs the ultraviolet radiation emitted by the scintillatorand converts it into long-wave radiation adapted to the spectralsensitivity of the photodiode. As a result, overlapping of the emissionspectrum of the scintillator and the sensitivity spectrum of thephotodiode is maximum and the photodiode can operate with a maximumquantum efficiency.

The invention also relates to a device for the detection of inputradiation which includes at least one acquisition element whichcomprises a sensor with a Pr³⁺-activated scintillator for converting theinput radiation into ultraviolet radiation and a photodiode whichconverts an optical signal into an electrical signal.

Embodiments of the invention will be described in detail hereinafterwith reference to the accompanying Figures. Therein:

FIG. 1 is a diagrammatic representation of the construction of a devicefor generating images and/or projections by means of the PET method,

FIG. 2 shows the excitation and emission spectrum of CaLi₂SiO₄:Pr, Na,

FIG. 3 shows the excitation and emission spectrum of LuPO₄:Pr,

FIG. 4 shows the excitation and emission spectrum of Lu₂SiO₅:Pr, and

FIG. 5 shows the excitation and emission spectrum of LaPO₄:Pr.

FIG. 6 shows a color converter that includes a polymer light guide and aseparate layer.

A device for generating images and/or projections can operate withvarious imaging methods. Preferably, the device is arranged to carry outthe PET method or the SPECT method as the imaging method, or to carryout the imaging method by means of X-rays.

In conformity with the PET method, a metabolic preparation marked withgiven, unstable nuclides is injected into a patient; this preparation istaken up in a tissue-specific or function-specific manner. Theradionuclides used decay, giving rise to two γ quanta in differentsuccessive processes in the vicinity of the location of decay; these twoquanta fly in exactly opposite directions and leave the patient so as tobe detected by the sensors which are arranged in the form of a ringaround the patient in the device for the detection of input radiation.During their travel from the location of origin to the location wherethey emerge from the patient the γ quanta traverse further tissue of thepatient; this tissue can absorb the γ quanta more or less in dependenceon the type of tissue. Generally speaking, the γ quanta are attenuatedin a tissue-specific manner. All detected γ quanta together form a setof projections of the image wherefrom a slice image or volume image isreconstructed in known manner during a subsequent reconstructionoperation. The PET method yields functional images of the object.

The two γ quanta have the same energy of 511 keV. The detection of the γquanta is performed by means of scintillators in the sensor of thedevice for the detection of input radiation.

FIG. 1 is a diagrammatic representation of a device for forming a sliceimage by means of the PET method. The patient, or the object to beexamined 1, is arranged within a ring-like device 2 for the detection ofinput radiation which consists of individual detection or acquisitionelements 3. The plane defined by the ring intersects the object 1, forexample, in the plane of intersection 4. The described decay processtakes place at the location 5 where two γ quanta leave the object 1 inopposite directions along the double arrow 6. The individual sensorelements 3 of the device 2 for the detection of input radiation areconnected, via leads 7, to a data processing unit 8 which evaluates thesignals from the detection elements 3. The slice image formed issubsequently displayed on a display device 8.

A high temporal resolution is very important for the PET method, becausethe entire process from the emission of the positron until the detectionof the γ quanta takes place within a few nanoseconds. Because of theshort decay time τ, the integration time during the determination of theintensity of the input radiation can be reduced, so that the image ratecan be significantly increased. As a result of the increased image rate,a higher dose of nuclides can be administered to the patient so as toreduce the overall examination time.

Furthermore, for the PET method it is also important to determine theenergy of the γ quanta in order to ensure that they have not left theiroriginal trajectory due to scattering processes. Such γ quanta have anenergy value which is lower than 511 keV. The energy value of the γquanta is determined from the level of the detected signal. In thisrespect it is important that the scintillator is in the normal stateagain and not in an excited state still, as otherwise an energy value of511 keV will be unduly determined, despite the fact that the γ quantumwas scattered and has a lower energy value.

The SPECT method also is a nuclear medical examination method. Thenuclides used for the SPECT method originate from natural decay andemitted γ quanta with energies of 141 keV (^(99m)Tc) at 159 keV (¹²³I).Like in the PET method, the emitted γ quanta are detected by the devicefor the detection of input radiation and their signal is amplified.However, a collimator is arranged in front of each detection element.The collimator serves as an objective and consists of a lead plate whichis provided with bores which are arranged in parallel or in a convergingfashion. γ quanta which are incident at an angle are absorbed by thecollimator, thus enabling the three-dimensional imaging.

In the case of X-ray computer-tomography the device for generatingimages or projections includes an X-ray tube which emits a fine, usuallyfan-shaped beam and moves in a circle around the longitudinal axis ofthe object to be examined. After having traversed the object, the X-raystransmitted by the tissue of the object are captured again by anoppositely situated device for the detection of input radiation. Forsmall fields of a format of approximately 1.5×1.5 mm the computercalculates the difference between the emitted energy and the receivedintensity of the X-ray beam and forms from the different values a greyscale image which is displayed on a display device. The grey valuescorrespond to the respective relative density of the tissue.

Like in the PET method, according to the latter two methods theintegration time during the detection of the intensity of the inputradiation can again be reduced as a result of the short decay time τ, sothat the image rate for the generating of images and/or projections canbe significantly increased. As a result of the increased image rate, theoccurrence of artifacts, such as shadow images, is reduced and also theperiod of time required to carry out the imaging method.

The device for the detection of input radiation may be composed of aplurality of acquisition elements, each acquisition element comprising asensor for converting the input radiation into UV radiation and aphotodiode. It is particularly advantageous when each acquisitionelement comprises an array of photodiodes. The sensor and the photodiodearray are customarily formed as a respective layer and combined so as toform a system of layers.

The sensor layer constitutes the entrance screen for the input radiationwhich is preferably formed by X-rays or γ rays. Underneath the sensorlayer there is arranged the layer with the photodiodes. Electricalcontact leads extend from the layer with the photodiodes to theelectronic read-out circuitry.

The essential element of the sensor is formed by a layer of aPr³⁺-activated scintillator which converts the input radiation into UVradiation. It is very advantageous when the sensor comprises a layer ofLaPO₄:Pr, LuF₃:Pr, LuCl₃:Pr, LuBr₃:Pr, (Lu_(1-x)Y_(x)):Pr, where 0≦x ≦1,(Lu_(2-x)Y_(x))SiO₅:Pr, where 0≦x≦1, (Lu_(1-x)Y_(x))Si₂O₇:Pr, where0≦x≦1, (Lu_(1-x)Y_(x))BO₃:Pr, where 0≦x≦1, :Pr_(y)Na_(y), where0.001≦y≦0.2. Such Pr³⁺-activated materials emit UV radiation afterexcitation and have short decay times.

TABLE 1 Emission bands and decay times of selected Pr³⁺-activated andCe³⁺-activated scintillators. Scintillator Emission wavelength λ_(max)[nm] Decay time τ [ns] LuPO₄:Pr 235 9 Lu₂SiO₅:Pr 270 16 LuPO₄:Ce 350 24Lu₂SiO₅:Ce 420 40The shorter decay time τ of the Pr³⁺-activated scintillators, that is,in comparison with the corresponding Ce³⁺-activated scintillators, canbe derived from the ratio τ≈1/λ_(max) ². The 5d levels of the free Pr³⁺cation are situated approximately 62000 cm⁻¹ above the 4f levels,whereas this energy difference amounts to only 50000 cm⁻¹ in the freeCe³⁺ cation. This larger energy difference is also responsible for thefact that the emission bands of Pr³⁺-activated scintillators,originating from 5d-4f transitions, are of higher energy than theemission bands of corresponding Ce³⁺-activated scintillators. Therefore,their main emission bands of the Pr³⁺-activated scintillators inaccordance with the invention are situated in the range from 220 to 350nm, so in the range of UV radiation, whereas the emission bands ofCe³⁺-activated scintillators are of a significantly longer wavelength.

The sensor layer with the Pr³⁺-activated scintillator is customarilyformed by pressing. A powder of the appropriate Pr³⁺-activatedscintillator is formed first; this powder is subsequently converted intoa crystal layer by means of a pressing method, for example,high-pressure cold pressing or high-temperature isostatic pressing. Thesize of such pressed crystals is in the range of from a few millimetersto a few centimeters. A sensor layer of this kind is bonded to the layerwith the photodiodes.

The powdery scintillator material can be formed from the startingcompounds by a solid-state reaction as well as by reactions in anaqueous or aqueous-alcoholic solution. During the latter reaction therelevant metal salts and/or metal oxides are possibly dissolved orsuspended with a combination of the anion of the scintillator in wateror a water-alcohol mixture so as to be reacted.

In the photodiodes the UV radiation is converted into electricalsignals. Because of the emission in the range of UV radiation, onlyUV-sensitive photodiodes can be used in the device for the detection ofinput radiation. For example, photodiodes with Cs₃Sb photocathodes,bialkaline photocathodes or multialkaline photocathodes can be used.Furthermore, use can also be made of photodiodes based on Si, GaN orAlGaN.

In order to enlarge the choice of suitable photodiodes, a colorconverter 10 which converts UV radiation into radiation of longerwavelength can be arranged between the sensor layer and the layer withthe photodiodes. To this end, the color converter 10 contains a luminoussubstance which can be excited by UV radiation. The color converter 10thus converts the UV radiation emitted by the sensor into radiationhaving a wavelength range which corresponds to the maximum of thespectral sensitivity of the photodiode used. The radiation of longerwavelength may comprise colored light or infrared radiation.

For the luminous substances use can notably be made of organic luminoussubstances with a high photoluminescence quantum efficiency and a shortdecay time τ. Particularly attractive luminous substances are Coumarinssuch as Coumarin 1 (λ_(max)=430 nm) or Courmarin 120 (α_(max)=442 nm) orLumogen dyes such as Lumogen F Violet 570 marketed by BASF. The decaytimes τ of these luminous substances are less than 10 ns.

Turning briefly to FIG. 6, the color converter 10 may comprise, forexample, a polymer light guide 11 which is doped with the luminoussubstance. The polymer light guide 11 may contain, for example,polymethylmethacrylate, polystyrol, polytetrafluoroethylene,polycarbonate, polyimide or polyvinylchloride. Alternatively, the colorconverter 10 may comprise two components, that is, the polymer lightguide 11 and a separate layer 12 with the luminous substance. In thisembodiment the polymer light guide 11 adjoins the sensor layer of theacquisition element; the separate layer 12 with the luminous substanceadjoins the polymer light guide 11 and is adjoined itself by the layerwith the photodiodes.

EXAMPLE 1

For the production of Ca_(0.98)Li₂SiO₄:Pr_(0.01),Na_(0.01), 25.0 g (278mmol) Li₂SiO₃, 147 mg (1.39 mmol) Na₂CO₃, 27.3 g (272 mmol) CaCO₃ and1.21 g (2.78 mmol) Pr(NO₃)₃.6H₂O were mixed in demineralized H₂O andsuspended. The water was removed by distillation and the residueobtained was dried. Subsequently, the reaction product was annealed oncein air at 700° C. for two hours and annealed twice in a CO atmosphere at850° C. for each time 12 hours. The scintillator powder was rinsedseveral times with water and ethanol, dried and ground on a rollerdevice for several hours. The resultant scintillator powder had anaverage particle size of 3 μm. The excitation and emission spectrum ofthis scintillator is shown in FIG. 2. The decay time τ ofCa_(0.98)Li₂SiO₄:Pr_(0.01),Na_(0.01) amounts to 16 ns.

EXAMPLE 2

40.0 g (101 mmol) Lu₂O₃ and 883 mg (2.03 mmol) Pr(NO₃)₃.6H₂O weresuspended in 200 ml ethanol in order to produce LuPO₄:Pr. 25.8 g (223mmol) 85% phosphoric acid was added slowly while stirring. Thesuspension was stirred for twenty-four hours and subsequentlyconcentrated in a rotary evaporator. The residue was dried at 100° C.,mortarized and mixed with 500 mg NH₄Cl. Subsequently, the scintillatorpowder was calcinated twice for 2 hours at 1250° C. in a CO atmosphereand subsequently mortarized each time again. Finally, the scintillatorpowder was calcinated once more in air at 1250° C. for 1 hour. Theexcitation and emission spectrum of this scintillator is shown in FIG.3. The decay time c of LuPO₄:Pr amounts to 9 ns.

EXAMPLE 3

10.0 g (25.1 mmol) Lu₂O₃, 1.51 g (25.1 mmol) and 86.0 mg (84.1 μmol)Pr₆O₁₁ were suspended in 200 ml ethanol in order to produce Lu₂SiO₅:Pr.The suspension was treated in an ultrasound bath for ten minutes andsubsequently concentrated in a rotary evaporator. The residue obtainedwas dried at 100° C., mortarized and mixed with 500 mg CsF.Subsequently, calcination took place in a CO atmosphere for 6 hours at1350° C. and the scintillator powder obtained was subsequentlymortarized. Finally, the scintillator powder was rinsed in 500 ml waterfor two hours, sucked off and dried at 100° C. The excitation andemission spectrum of this scintillator is shown in FIG. 5. The decaytime τ of Lu₂SiO₅:Pr amounts to 16 ns.

EXAMPLE 4

91.4 g (282 mmol) La₂O₃ and 883 mg (2.48 mmol) PrCl₃.6H₂O were suspendedin 200 ml water in order to manufacture LaPO₄:Pr. 69.0 g (598 mmol) 85%phosphoric acid was added slowly while stirring. The suspension wasstirred for 24 hours and subsequently concentrated in a rotaryevaporator. The residue obtained was dried at 100° C., mortarized andmixed with 1.2 g LiF. The scintillator was calcinated in a nitrogenatmosphere for 2 hours at 1000° C. The scintillator powder obtained wasrinsed in diluted HNO₃ for 6 hours at 60° C. Finally, the scintillatorwas sucked off, rinsed acid free with water and dried at 100° C. FIG. 6shows the excitation and emission spectrum of this scintillator. Thedecay time τ of LaPO₄:Pr amounts to 11 ns.

1. A device for generating images and/or projections by means of an imaging method, which device includes a device for the detection of input radiation which includes at least one acquisition element which comprises a sensor with a Pr³⁺-activated scintillator for converting the input radiation into UV radiation, a color converter which contains a luminous substance for converting the UV radiation to an optical signal, and a photodiode which converts an optical signal into an electrical signal; wherein the Pr³⁺-activated scintillator is chosen from the group LuF₃: Pr, LuCl₃:Pr, and LuBr₃:Pr.
 2. A device for generating images and/or projections as claimed in claim 1, wherein the device is arranged to carry out the PET method as the imaging method.
 3. A device for generating images and/or projections as claimed in claim 1, wherein the device is arranged to carry out the SPECT method as the imaging method.
 4. A device for generating images and/or projections as claimed in claim 1, wherein the device is arranged to carry out the imaging method by means of X-rays.
 5. A device for generating images and/or projections as claimed in claim 1, wherein the color converter comprises a polymer light guide which is doped with the luminous substance that is excited by UV radiation.
 6. A device for generating images and/or projections as claimed in claim 1, wherein the color converter comprises a polymer light guide and a separate layer with the luminous substance that is excited by UV radiation.
 7. A device for generating images and/or projections as claimed in claim 1, wherein the luminous substance includes an organic material.
 8. A device for the detection of input radiation which includes at least one acquisition element which comprises a sensor with a Pr³⁺-activated scintillator for converting the input radiation into UV radiation, a color converter that converts UV radiation to an optical signal, and a photodiode which converts the optical signal into an electrical signal, wherein the color converter includes a polymer light guide and the color converter is doped with a Courmarin based substance.
 9. A device for detecting input radiation as claimed in claim 8, wherein the color converter contains a luminous substance which can be excited by UV radiation, wherein the color converter is arranged between the sensor and the photodiode.
 10. A device for detecting input radiation as claimed in claim 8, wherein the acquisition element comprises an array of photodiodes, and further wherein the array of photodiodes forms a first layer and the sensor forms a second layer, wherein the first and second layers are combined to form a system of layers.
 11. A device for detecting input radiation as claimed in claim 8, wherein a decay time of the scintillator is approximately 9 ns.
 12. A device for detecting input radiation as claimed in claim 8, wherein a decay time of the scintillator is approximately 16 ns.
 13. A device for detecting input radiation as claimed in claim 8, wherein the Pr³⁺-activated scintillator is Ca_(1-2y)Li₂SiO₄:Pr_(y)Na_(y), where 0.001≦y≦0.2.
 14. A device for detecting input radiation as claimed in claim 8, wherein the Pr³⁺-activated scintillator is LaPO₄:Pr.
 15. A device for detecting input radiation as claimed in claim 8, wherein the Pr³⁺-activated scintillator is one of LuCl₃:Pr, LuBr₃:Pr, (Lu_(2-x)Y_(x))SiO₅:Pr, where 0 ≦x≦1, and (Lu_(1-x)Y_(x))Si₂O₇:Pr, where 0≦x≦1.
 16. An imaging method, comprising: receiving one of an X-ray and a γ quantum at a Pr³⁺-activated scintillator, wherein the Pr³⁺-activated scintillator is one of LuCl₃:Pr, LuBr₃:Pr; receiving UV radiation emitted form the scintillator at a color converter in response to receipt of the one of the XC-ray and the γ quantum; receiving a light signal emitted from the color converter at a photodiode; generating an electrical signal in response to receipt of the light signal; and generating an image based at least in part upon the generated electrical signal.
 17. The imaging method of claim 16, wherein the color converter includes a polymer light guide that is doped with a luminous substance.
 18. The imaging method of claim 16, wherein the color converter includes a polymer light guide and a separate layer with a luminous substance.
 19. The device for detecting input radiation as claimed in claim 16, wherein the color convener is doped with a Courmarin based substance. 